In vertebrate animals, the heart is a hollow, muscular organ having four pumping chambers: the left and right atria and the left and right ventricles, each being provided with its own one-way valve. Thus, there are four heart valves: the mitral valve and the tricuspid valve, which are called atrio-ventricular valves, and the pulmonary valve and the aortic valve, which are called ventriculo-arterial valves. In one cycle of cardiac contraction, a valve opens to let blood flow from one side to the other, and then it closes to prevent backflow in the other direction. Thus, in the diastolic phase, the atrio-ventricular valves open to enable filling of the ventricles while the ventriculo-arterial valves remain closed. Conversely, in the systolic phase (ventricular contraction) of the cardiac cycle, the mitral and tricuspid valves close, while the pulmonary and aortic valves open to enable ejection of blood downstream of the ventricles.
Heart valve disease or dysfunction can, in severe cases, substantially restrict day to day activities and shorten the lives of patients. The primary procedure is valve repair or replacement surgery. In more than 20 percent of cardiac surgeries, heart valve disease represents the principal reason for a type of cardiac operation known as “open heart” surgery. These operations are associated with significant morbidity and mortality according to multiple risk factors related to age. New minimally-invasive procedures have been proposed to improve overall success. They involve inserting a valve made from animal tissue into the body using catheters and placing it inside the original diseased valve. However, both valve repair and replacement techniques remain particularly expensive and are associated with important risks for the patients.
It is therefore important to be able to quantify accurately the severity of valve disease, identify the right diagnosis and provide the proper treatments to patients. In addition, following a heart valve procedure, it is also fundamental to assess and monitor the physiologic performance of the new or repaired valve.
Heart valve pathologies may include a defect in closing or opening of a valve, or a combination of these two dysfunctions. Epidemiologically, a defect in opening of the aortic valve remains one of the most frequent anomalies. The three usual causes of aortic valve stenosis remain, in order of importance: calcified degeneration of the valve attributable to age, but certainly accentuated by hypertension; a congenital anomaly called bicuspid aortic valve, in which the valve possesses only two rather than three cusps (a cusp may also be referred to as a leaf or leaflet); and acute rheumatic fever (a particular bacterial infection at an early age, with subsequent excess scarring).
Diagnosis of aortic valve stenosis relies, first and foremost, on the appearance of symptoms in the patient, namely, breathlessness and chest pain on exertion, as well as loss of consciousness. Diagnosis is made through physical examination of the patient revealing a characteristic systolic murmur on auscultation, and a slower than normal rise of the pulse. Following that, transthoracic echography may confirm the clinical impression, e.g. by measuring an abnormal thickening of the layers of the heart valves, presence of calcifications, as well as restriction of the movement of the opening of the valve (Braunwald's Heart Disease: A Textbook of Cardiovascular Medicine, Authors: Peter Libby, Robert O. Bonow, Douglas L. Mann and Douglas P. Zipes, pages 267-268). The Doppler effect can be used to measure the velocity of blood traversing through the valve. This type of examination (two dimensional imaging and Doppler) allows for evaluation of the maximal velocity of the transvalvular blood flow, the aortic valve surface area and the surface area of the left ventricle outflow tract (LVOT). Accordingly, a severe aortic stenosis is defined by:
a) maximal flow velocity of >4.5 m/sec;
b) a mean transvalvular pressure gradient>50 mmHg;
c) a ratio of the surface of the LVOT/aortic valve×TVI (Time Velocity Interval)<0.25;
or
d) an estimated valve surface area of <0.75 cm2 (Braunwald's Heart Disease: A Textbook of Cardiovascular Medicine, Authors: Peter Libby, Robert O. Bonow, Douglas L. Mann and Douglas P. Zipes, page 269).
Even though echo-cardiography remains the most frequently used method for confirmation of diagnosis, it is limited by the echogenicity of the patient, valvular calcifications and sub-valvular calcifications, as well as the presence of concomitant mitral valve pathology. The results of echo-cardiography are also strongly operator-dependent. Consequently, it would be preferable to have a diagnostic method that relies on an in situ measurement of the transvalvular pressure gradient, made directly via cardiac catheterization into the interior of the heart in the region of the valve concerned, and surrounding regions.
In situ measurements of pressure or flow can be made within the human body, by minimally-invasive techniques, using a pressure sensing catheter or a special guidewire equipped with an integrated pressure sensor (see, for example, Grossman's cardiac catheterization, angiography, and intervention, Authors: Donald S. Baim and William Grossman, pages 647-653).
Classically, this measurement of a transvalvular pressure difference, or pressure gradient, is made by two pressure sensing catheters positioned upstream and downstream, respectively, of the valve of interest, e.g. the aortic valve. To do this, the transeptal approach uses a Brockenbrough needle, and a Mullins type catheter to permit access, via the femoral vein, to the left atrium, by perforating the inter-atrial septum, then traversing the mitral valve, to place the tip of the catheter in the left ventricular cavity. A second catheter is introduced through the common femoral artery, to be positioned in the ascending aorta, just above the cusps of the aortic valve. Simultaneous measurements of the pressure upstream and downstream of the aortic valve can thus be obtained.
An alternative to this method involves placing a first arterial catheter through the aortic valve in the left ventricle and then introducing a second arterial catheter, and positioning it above the aortic valve in the ascending aorta. However, in this technique there is an exaggeration of the transvalvular pressure gradient since one of the catheters rests through the valve.
In another method, one can also simply measure the transvalvular gradient with only one arterial catheter, by a “pull back” method, i.e. inserting the catheter, crossing the valve into the ventricle and once the ventricular pressure measurement has been made, quickly withdrawing from the left ventricle into the ascending aorta, and subsequently measuring the aortic pressure. This last technique is clearly less reliable since, firstly, the measurements of the aortic and ventricular pressures are not simultaneous and secondly, the withdrawal of the catheter frequently involves a benign transitory cardiac arrhythmia, which distorts the pressure curves.
A variant of the latter approach was reported in an experimental study (Feasibility of a Pressure Wire and Single Arterial Puncture for Assessing Aortic Valve Area in Patients with Aortic Stenosis, J. H. Bae et al., J. Invasive Cardiol., 2006 August, 18(8), pp. 359-62). A pressure sensing wire, inserted through a guiding catheter, was used to measure pressure in the left ventricle using the pressure sensing wire, while simultaneously measuring pressure in the aorta using the guiding catheter. In practice, the technique is used infrequently. Firstly, it is not ideal to use different types of devices for comparing the two pressures. Also, in practice, the pressure sensing wire described is used for measuring pressures within small blood vessels, such as coronary arteries, and is therefore of small diameter and very flexible. It is therefore too limp and fragile for reliably positioning it for pressure measurements within the heart, where higher blood flows and significant turbulence in the flow tends to cause movement of the sensor at the end of the wire.
To calculate the cardiac flow, thermodilution by a Swan-Ganz catheter or the method of Fick are commonly used (Grossman's cardiac catheterization, angiography, and intervention, Authors: Donald S. Baim and William Grossman, pages 150-156).
Besides diagnosing a heart valve condition, measurements of a blood pressure gradient in blood vessels may be used to diagnose and treat patients with multi-site vessel disease. In order to quantify lesion severity in a diffusely affected vessel, pressure measurements are made at several locations along the vessel. This is currently done by withdrawing a pressure sensor equipped guidewire along a length of the vessel from a distal to a proximal position very slowly during a steady-state maximum induced hyperaemia. This diagnostic shows the location and severity of lesions but accuracy is compromised by the sequential nature of the data.
In view of limitations, such as limited accuracy, of the above-mentioned apparatus and techniques, there is a need for improved or alternative systems, apparatus and methods of operation for directly measuring and monitoring blood pressure gradient in real time, more accurately and reliably than is now possible, using minimally-invasive techniques.
A pressure sensing catheter is effectively a fluid-filled catheter: a pressure at the distal end, positioned in the region of interest, is measured by monitoring the fluid pressure in the catheter at the proximal end. A pressure sensing catheter for sensing pressure within the heart is typically 6 to 8 French in outer diameter (0.078″ to 0.104″), in order to maintain enough rigidity and robustness. Typically, a pressure sensing guidewire equipped with an electrical pressure sensor can be made with a smaller diameter. This is advantageous for applications such as transvalvular pressure measurements or for measurements in small blood vessels such as coronary vessels.
One type of commercially available sensor equipped guidewire, PressureWire Certus from St. Jude Medical, uses a Micro-Electro-Mechanical-Systems (MEMS) device that includes a piezoresistor and diaphragm, e.g. as described in U.S. Pat. Nos. 5,343,514 and 6,615,667 to Smith (Radi Medical Systems AB) entitled “Combined flow, pressure and temperature sensor”. Deformation of the diaphragm, caused by a pressure change, is read using resistance values. Other similar systems using MEMS technologies monitor the capacitance value between a fixed plate and the diaphragm to evaluate the deformation of the diaphragm due to pressure changes.
As explained above, available single pressure sensor guidewires can measure pressure at only one point at a time, and to measure a pressure gradient, the guidewire sensor must be moved through a region of interest, such as through a heart valve, or other vascular region, to measure pressure sequentially at several different points.
A problem with guidewires equipped with sensors based on electrical signals is that multiple, long electrical connections to each sensor must be provided. The length of a guidewire may be more than 1 meter. Use of microelectronics and long electrical wires, particularly when used in humid biological conditions, tends to cause reliability issues with measurement of small electrical signals, e.g. from parasitic capacitances, noise and electromagnetic interference (EMI), and limits the ability to integrate multiple electrical sensors within a guidewire to measure pressure gradient and flow. Furthermore, there may be significant risks involved with the use of microelectronics and electrical connections in vivo, particularly in the region of the heart, where electrical activity may disrupt normal heart function.
The electronic drift of MEMS sensors integrated into guidewires remains a limitation. For example, in one study, it was reported that measured pressures dropped >5 mmHg/hour, due to drift, therefore causing pressure gradient over-estimation (Coronary Pressure, Authors: Nico Pijls and Bernard de Bruyne, pages 125-127).
Additionally, the guidewire is fabricated to provide the required flexibility and torque characteristics to enable the guidewire to be steered and positioned. Thus, the guidewire usually includes torque steering components comprising a central core-wire or mandrel, and external coil, e.g. a fine spiral metal coil, and a J-shaped tip (pre-shaped or manually shaped).
A guidewire used for cardiology may typically have a gauge of between 0.89 mm (0.035″) and 0.25 mm (0.010″) for introduction into small blood vessels. Note: catheter gauge may also be specified in French units: 1 French=0.333 mm diameter (0.013″). It will be appreciated that there is a limit to the number of electrical wires, sensors and steering components that can physically fit within the required diameter guidewire. Even if larger guidewires could be inserted, they would tend to interfere with normal operation of a heart valve and distort measurements, so it is desirable that the guidewire is as small gauge as possible. This presents a number of challenges in providing a guidewire with more than one electrical sensor.
In addition, MEMS sensors along with their long electrical connections significantly increase the complexity of manufacturing assembly processes of guidewires using electrical sensors, and therefore significantly increasing their manufacturing costs. Typically, guidewires for medical use are fabricated to be disposable (i.e. for single use only) and are significantly expensive.
To provide multi-sensor capability with a single electrical connection, U.S. Pat. No. 6,615,667 discloses a single combined flow, pressure and temperature MEMS sensor, but again, pressure can be measured at only one point.
To avoid the need for wired electrical connections entirely, optical pressure sensors are also known which are optically coupled to the control unit by optical fibers. However, another challenge for medical applications, as described above, is that to measure pressure gradients, pressure sensors are required having sufficient sensitivity to detect small pressure differences reliably within the region of interest. Some available optical sensors are either too large to allow for multiple sensors to be accommodated in a small gauge device, and/or they do not have sufficient sensitivity.
U.S. Pat. No. 4,735,212 to Cohen (Cordis Corporation) entitled “Multiple site fiber-optic pressure transducer” and U.S. Pat. No. 4,543,961 to Brown (Cordis Corporation) entitled “Data transmission system” discloses early designs for integrating several miniaturized pressure transducers or sensors arranged in a relatively large, i.e. 1.5 mm (0.060″) single fiber device. These designs are quite complex and would appear to be a challenge to fabricate consistently. More significantly, the sensor elements will be sensitive to stresses when the fiber is bent or twisted, such that it would be difficult to discriminate fiber stresses from actual pressure readings. Thus, even if they could be manufactured with sufficiently small diameters, these and similar configurations would not be suitable for intravascular or intravalvular use which necessitate bending of the fiber in the region of the sensors.
Another known type of single point optical pressure sensor is a Micro-Opto-Mechanical Systems (MOMS) device that comprises a Fabry-Perot optical cavity where one of the two mirrors is a diaphragm. Low-coherence light is sent to the cavity via an optical fiber. Diaphragm motions are measured from spectral changes of the reflected light. Miniaturized pressure sensors of this type are described, for example, in U.S. Pat. No. 6,684,657 to Donlagic et al. (Fiso Technologies Inc.) entitled “Single Piece Fabry-Perot Optical Sensor and Method of Manufacturing the Same”, and also in U.S. Pat. No. 7,689,071 to Belleville et al. (Opsens Inc.) entitled “Fiber-optic pressure sensor for catheter use”. The use of this type of sensor for use in cardiovascular applications is relatively recent.
In summary, existing guidewire apparatus, using various types of sensors, are available for single point pressure measurements, for example, from St. Jude Medical and Volcano Corporation. However, apparatus is not currently known or available to cardiologists for directly measuring in situ blood pressure gradients simply and quickly, particularly transvalvular pressure gradients, where a catheter with a diameter of 0.89 mm (0.035″) and preferably 0.46 mm (0.018″) or less is needed to minimize disruption to normal heart valve activity and over estimation of transvalvular gradient. It would also be desirable to enable measurements for simultaneous determination of cardiac output and valvular area.
Thus, there is a need for improved or alternative systems, apparatus and methods for direct measurement and monitoring of blood pressure, pressure gradients and/or flow within the heart and the vascular system, and in particular, for measurement of transvalvular pressure gradients and flow velocity.